Laser system and processing conditions for manufacturing bioabsorbable stents

ABSTRACT

The present invention involves laser machining polymer substrates to form a stent with laser parameters that minimize damage to the substrate in a surface region adjacent to the machined edge surface. The wavelength and pulse width are selected for this unique application and they can be controlled to minimize the surface modifications (such as voids, cracks which are induced by the laser-material interaction) which contribute to the variation in mechanical properties with distance from the edge surface, bulk mechanical properties, or a combination thereof.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to laser machining tubing to form stents.

2. Description of the State of the Art

This invention relates to laser machining of devices such as stents.Laser machining refers to removal of material accomplished through laserand target material interactions. Generally speaking, these processesinclude laser drilling, laser cutting, and laser grooving, marking orscribing. Laser machining processes transport photon energy into atarget material in the form of thermal energy or photochemical energy.Material is removed by melting and blow away, or by directvaporization/ablation.

When a substrate is laser machined energy is transferred into thesubstrate. As a result, a region beyond the cutting edge is modified bythe energy, which affect the properties in this region. In general, thechanges in properties are adverse to the proper functioning of a devicethat is being manufactured. Therefore, it is generally desirable toreduce or eliminate energy transfer beyond removed material, thusreducing or eliminating the extent of modification and size of theregion affected.

One of the many medical applications for laser machining includesfabrication of radially expandable endoprostheses, which are adapted tobe implanted in a bodily lumen. An “endoprosthesis” corresponds to anartificial device that is placed inside the body. A “lumen” refers to acavity of a tubular organ such as a blood vessel.

A stent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices, which function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of the diameter of a bodily passage ororifice. In such treatments, stents reinforce body vessels and preventrestenosis following angioplasty in the vascular system. “Restenosis”refers to the reoccurrence of stenosis in a blood vessel or heart valveafter it has been treated (as by balloon angioplasty, stenting, orvalvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through a bodily lumen to a region, such as alesion, in a vessel that requires treatment. “Deployment” corresponds tothe expanding of the stent within the lumen at the treatment region.Delivery and deployment of a stent are accomplished by positioning thestent about one end of a catheter, inserting the end of the catheterthrough the skin into a bodily lumen, advancing the catheter in thebodily lumen to a desired treatment location, expanding the stent at thetreatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon. The stent is thenexpanded by inflating the balloon. The balloon may then be deflated andthe catheter withdrawn. In the case of a self-expanding stent, the stentmay be secured to the catheter via a retractable sheath or a sock. Whenthe stent is in a desired bodily location, the sheath may be withdrawnwhich allows the stent to self-expand.

The stent must be able to satisfy a number of mechanical requirements.First, the stent must be capable of withstanding the structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel. Therefore, the stent must possess adequate radialstrength and rigidity. Radial strength is the ability of a stent toresist radial compressive forces. Once expanded, the stent mustadequately maintain its size and shape throughout its service lifedespite the various forces that may come to bear on it, including thecyclic loading induced by the beating heart. For example, a radiallydirected force may tend to cause a stent to recoil inward. Generally, itis desirable to minimize recoil.

In addition, the stent must possess sufficient flexibility to allow forcrimping, expansion, and cyclic loading. The stent should havesufficient resistance to fracture so that stent performance is notadversely affected during the crimping, expansion, and cycling loading.

Finally, the stent must be biocompatible so as not to trigger anyadverse vascular responses.

The structure of a stent is typically composed of scaffolding thatincludes a pattern or network of interconnecting structural elementsoften referred to in the art as struts or bar arms. The scaffolding canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. The scaffolding is designed so that the stent can beradially compressed (to allow crimping) and radially expanded (to allowdeployment).

Stents have been made of many materials such as metals and polymers,including biodegradable polymeric materials. Biodegradable stents aredesirable in many treatment applications in which the presence of astent in a body may be necessary for a limited period of time until itsintended function of, for example, achieving and maintaining vascularpatency and/or drug delivery is accomplished.

Stents can be fabricated by forming patterns on tubes or sheets usinglaser machining Even though the basic laser-material interaction issimilar, there are certain aspects among types of materials (such asmetals, plastics, glasses, and ceramics), i.e. different absorptioncharacteristics. To produce the desired results, it is critical inchoosing a suitable wavelength. Once a suitable wavelength is selected,it is the combination of pulse energy and pulse duration that define theoptimal process condition for the type of material. The properties ofbiodegradable polymers like PLLA and PLGA tend to be very sensitive toenergy transfer such as that from laser machining There are greatefforts needed in understanding the laser parameters and laser—materialinteraction to help select a laser system and define processingparameters enabling faster laser machining of biodegradable stents whichminimize the adverse effects on the properties.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a method of lasermachining a substrate to form a stent, comprising: providing athin-walled polymer substrate; laser machining the thin-walled polymersubstrate with a laser beam with a pulse width and wavelength that cancut through the wall of the substrate to form structural elements havinga machined edge surface, wherein the laser beam modifies the substratein a surface region adjacent to the machined edge surface in much lessdegree, wherein the modifications include voids, cracks that contributeto the variation in modulus of the polymer with distance from the edgesurface, or a combination thereof, and selecting the pulse width andwavelength so that the voids or cracks are present at no greater than adepth of 2 microns or the modulus converges at no greater than 4microns.

Additional embodiments of the present invention include a method oflaser machining a substrate to form a stent, comprising: providing athin-walled PLLA polymer substrate; laser machining the thin-walled PLLApolymer substrate with a laser beam to cut through the wall to formstructural elements having a machined edge surface, wherein the pulsewidth and wavelength of the laser beam are within the green range andthe pulse width is 1-10 ps.

Further embodiments of the present invention include a polymer stentbody, comprising: a plurality of interconnected structural elementsformed by laser machining a thin-walled PLLA polymer substrate with alaser beam that cuts through the wall to form the structural elements,wherein the structural elements have sidewalls corresponding to amachined edge surface, wherein a surface region adjacent to thesidewalls has damage caused by interaction of the laser beam with thesubstrate, wherein the damage comprises voids or cracks dispersed in thesurface region to a depth of 2 microns or less from the edge surface.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIG. 2 depicts a machine-controlled system for laser machining a tube.

FIG. 3 depicts a close-up axial view of a region where a laser beaminteracts with a tube.

FIG. 4A depicts a portion of a strut or structural element formed bylaser machining a substrate.

FIG. 4B depicts a cross-section of a portion of a strut normal to amachined edge surface.

FIGS. 5-8 show SEM images of the surface region adjacent to a lasermachined edge surface for combinations of pulse width and wavelength.

FIGS. 9-13 show the modulus vs. displacement into the surface of a lasermachined edge for combinations of pulse width and wavelength.

FIGS. 14-17 depict SEM images showing the sidewall surfaces of stentsmachined with different combinations of pulse widths and wavelengths.

DETAILED DESCRIPTION OF THE INVENTION

Embodiments of the present invention relate to methods of lasermachining of polymer substrates to manufacture a stent. Morespecifically, these embodiments relate to selecting and implementing alaser system and parameters that reduce or eliminate adverse effects ofthe laser on the polymer and which maximize the preservation of thefunctional properties of the material, such as the surface and bulkproperties of the polymer material.

In general, stents can have virtually any structural pattern that iscompatible with a bodily lumen in which it is implanted. Typically, astent is composed of a pattern or network of circumferential rings andlongitudinally extending interconnecting structural elements of strutsor bar arms. In general, the struts are arranged in patterns, which aredesigned to contact the lumen walls of a vessel and to maintain vascularpatency.

An exemplary structure of a stent is shown in FIG. 1. FIG. 1 depicts astent 10 which is made up of struts 12. Stent 10 has interconnectedcylindrical rings 14 connected by linking struts or links 16. Theembodiments disclosed herein are not limited to fabricating stents or tothe stent pattern illustrated in FIG. 1. The embodiments are easilyapplicable to other stent patterns and other devices. The variations inthe structure of patterns are virtually unlimited. The outer diameter ofa fabricated stent (prior to crimping and deployment) may be between0.2-5.0 mm. For coronary applications, a fabricated stent diameter is2.5-5 mm. The length of the stents may be between about 6-12 mm or moredepending on the application.

The present embodiments are particular relevant to laser machiningpolymer substrates to form stents, however, the methods may beapplicable to other materials such as metals and ceramics or compositematerials composed of combinations of polymer, metal, and ceramic.

Polymers can be biostable, bioabsorbable, biodegradable, or bioerodable.Biostable refers to polymers that are not biodegradable. The termsbiodegradable, bioabsorbable, and bioerodable, as well as degraded,eroded, and absorbed, are used interchangeably and refer to polymersthat are capable of being completely eroded or absorbed when exposed tobodily fluids such as blood and can be gradually resorbed, absorbed,and/or eliminated by the body. In addition, a medicated stent may befabricated by coating the surface of the stent with an active agent ordrug, or a biodegradable polymer carrier including an active agent ordrug. The drug coating is typically applied to the stent body orscaffolding after being formed by laser machining The coating istypically much thinner than the struts of the scaffolding, for example,the coating can be 1-5 microns in thickness while the struts aretypically greater than 100 microns in thickness, for example, 140-160microns thick.

An implantable medical device, such as a stent, can be fabricated bylaser machining a construct or substrate to form the device. Material isremoved from selected regions of the construct which results information of the structure of the device. In particular, a stent may befabricated by machining a thin-walled tubular member with a laser.Selected regions of the tubing may be removed by laser machining toobtain a stent with a desired pattern. Specifically, a laser beam can bescanned over the surface of a tubing or the tubing can be translated androtated under the beam resulting in removal of a trench or kerfextending all the way through a wall of the tubing. When a starting andending point of a kerf meet, the region surrounded by the kerf drops outor is removed by an assisting gas. FIG. 2 depicts an embodiment of aportion of a machine-controlled system for laser machining a tube. InFIG. 2, a tube 200 is disposed in a rotatable collet fixture 204 of amachine-controlled apparatus 208 for positioning tubing 200 relative toa laser 212. According to machine-encoded instructions, tube 200 isrotated and moved axially relative to laser 212 which is alsomachine-controlled. The laser selectively removes the material from thetubing resulting in a pattern cut into the tube. The tube is thereforecut into the discrete pattern of the finished stent.

FIG. 3 depicts a close-up view of a laser beam 408 interacting with atube 414. Laser beam 408 is focused by a focusing lens 338 on tube 414.Tube 414 is supported by a controlled rotary collet 337 at one end andan optional tube support pin 339 at another end. A coaxial gas jetassembly 340 directs a cold gas jet or stream 342 that exits through anozzle 344 that cools the machined surface as the beam cuts and ablatesa substrate. The gas stream also helps to remove debris from the kerfand cool the region near the beam. Gas inlet is shown by an arrow 354.Coaxial gas jet nozzle 344 is centered around a focused beam 352. Insome embodiments, the pressure of the supplied cooling gas is between 30and 150 psi. An exemplary flow rate of the cooling gas is between 2 and100 scfh. Exemplary cooling gases or process gases include helium,argon, nitrogen, oxygen, or a mixture of those gases.

Biodegradable polymers that may be suitable for stent scaffoldapplications include semi-crystalline polymers. In particular, theseinclude polymers with a glass transition temperature (Tg) above humanbody temperature, which is about 37° C. The polymer substrate can bemade in whole or in part from single or a combination of biodegradablepolymers including, but not limited to, poly(L-lactide) (PLLA),polymandelide (PM), poly(DL-lactide) (PDLLA), polyglycolide (PGA), andpoly(L-lactide-co-glycolide) (PLGA). For PLGA, it includes copolymercontaining different molar ratios of L-lactide to glycolide, such as,90:10, 75:25, 50:50, 25:75, and 10:90.

Several properties of a stent are essential for performing its functionincluding, radial strength and fracture resistance or elongation atbreak. For example, adequate fracture resistance is required and crucialsince the stent undergoes significant stress/strain in localized regionswhen the stent is crimped and deployed. The inside or concave regions20, illustrated in FIG. 1, of the bends in the stent pattern or crowns18 are subjected to high compressive stress and strain when the stent iscrimped, but the outside or convex regions 22 of the crowns 18 aresubjected to high compressive stress and strain when the stent isdeployed. Thus, the stent during crimping and deployment is highlysusceptible to cracking Such cracking can lead to a loss of radialstrength and potentially to premature and/or to a catastrophic failureof the stent. Therefore, it is essential that the mechanical propertiesof the polymeric stent can be maintained through the laser machiningprocess.

To provide the stent with the desired properties, an additional processstep can be introduced to enhance radial strength and fractureresistance to the pre-form, tubing, prior to the its lasing step. Forexample, fracture toughness can be greatly enhanced by controlling thesize of crystalline domains and by optimizing an optimalamorphous/crystalline ratio for a semi-crystalline polymer; radialstrength is also seen to be enhanced by preferential polymer chainalignment to the tube in the hoop direction. These desiredmicrostructural properties can be tailored through a heated radialexpansion step on the polymer tube above the Tg of the polymer. Forexample, for PLLA, a range 65-120° C. is preferred. On the contrary,localized heating to the substrate during laser cutting can result inmodification of the desired microstructural properties, or damage tolocalized regions, which can result loss or reduction of the advantagesprovided by the processing of the tube.

Laser beam machining is one of most advanced non-contact type machiningtechnology used in micro and nano-fabrication to fulfill the need inhandling advanced engineering materials, stringent design requirements,intricate shape and unusual size. The present invention relates tolasers having pulse widths in a range of a picosecond (=10⁻¹²)(“Picosecond” lasers) and having pulse widths in the range of afemtosecond (=10⁻¹⁵). “Pulse width” refers to the duration of an opticalpulse versus time. The duration can be defined in more than one way.Specifically, the pulse duration can be defined as the full width athalf maximum (FWHM) of the optical power versus time. Picosecond andfemtosecond lasers offer unique advantages for the removal of preciseamount of materials with minimum thermal damage to the surroundingmaterials. In general, the picosecond lasers have pulse widths less thanabout 10 ps and the femtosecond lasers have pulse widths between 10 and800 fs.

The two fundamental mechanisms involved in the laser ablation arebelieved to be photothermal and photochemical mechanism. In thephotothermal mechanism the material is ablated by melting andvaporizing, whereas in photochemical mechanism the photo energy of lightis used to break the chemical bonds of the polymer directly. Thechemical bonds between atoms and molecules of the substrate are brokenresulting in formation of gaseous species which are removed from thesubstrate.

Laser ablation of material from a substrate can occur by a thermalmechanism, a nonthermal mechanism, or a combination of both.Longer-pulse lasers, for example, remove material from a surfaceprincipally through a thermal mechanism. In a thermal mechanism, thelaser energy that is absorbed results in a temperature increase at andnear the absorption site and material is removed by conventional meltingor vaporization. The disadvantage of machining by this mechanism is theoccurrence of thermal damage of uncut substrate material. Such damageincludes melting and thermal diffusion into a region or zone of materialat the machining edge which results in modification of the properties ofthe substrate in the zone and cut quality problems.

Lasers with femtosecond pulse duration are of particular interest forablating material recently as the pulse duration is less than thetypical thermalization characteristic time (i.e., time to achievethermal equilibrium) of a few picoseconds. Due to a much smaller thermaldiffusion depth, it is considered as a completely or nearly completelynonthermal mechanism. A picosecond laser removes material mostly througha nonthermal mechanism, but also with some degree of thermal mechanismfor some materials that is enough to cause some thermal damage to asubstrate.

More specifically, the nonthermal mechanism involves optical breakdownin the target material which results in material removal. During opticalbreakdown of material, a very high free electron density, i.e., plasma,is produced through mechanisms such as multiphoton absorption andavalanche ionization. With optical breakdown, the target material isconverted from its initial solid-state directly into fully ionizedplasma on a time scale too short for thermal equilibrium to beestablished with a target material lattice. Therefore, there isnegligible heat conduction beyond the region removed. As a result, thereis negligible thermal stress or shock to the material beyondapproximately 1 micron from the laser machined surface.

However, it is believed that it is not known in the prior art whetherthe nonthermal or photochemical mechanism can cause damage to the uncutsubstrate that adversely effects stent performance, in particularfracture resistance. Specifically, it is not clear that opticalbreakdown propagates or penetrates beyond the laser machined edgesurface and causes damage within the substrate. It is also not known ifsuch damage can influence properties that effect stent performance.Furthermore, the existent relationship of laser parameters in ablatingpolymeric materials such as wavelength, pulse energy, and pulse width ofthe laser beam to such potential damage is not known.

We have learned from our studies that although laser ablating PLLAmaterial by a nonthermal mechanism by using a femtosecond laser hasnegligible thermal damage to the cut surface, the photochemical effectdoes cause damages (void and crack) under the surface region which isadjacent to the laser machined edge surface. These damages surely lowerthe stent performances. Therefore, it is critical in choosing a lasersystem and its processing parameters to fabricate polymer stents withminimum damages to the properties of stent such as fracture resistance.

The embodiments of the present invention include defining processparameters on a picosecond pulse laser system for making polymericstents with preserved mechanical properties, such as radial strength,elongation at break, or fracture resistance. The laser parameters thatare controlled include the pulse width and wavelength of the laserenergy. The use of such parameters in laser machining can minimizedamages arising from both thermal and nonthermal mechanisms. Theembodiments further include implementing a laser system and itsparameters on a laser machining used for making PLLA and PLGA-basedsubstrates have minimum damage.

The inventors recognized that the extent and depth of the damage causedby the photochemical mechanism can be controlled by laser parameterssuch as pulse width and wavelength. In general, the inventors haverecognized that in laser machining polymers there is a tradeoff betweendamage cause by a thermal and nonthermal mechanism. As discussed below,both thermal and nonthermal mechanisms (i.e., photochemical) causedamage to the substrate. Although the character of the damage isdifferent, both can adversely effect stent performance. The laserparameters (e.g., pulse width and wavelength) can be adjusted to reducethe photochemical effects, but with some increase in the thermalmechanism.

To illustrate damages caused by shorter laser pulses, PLLA stent sampleswere made by using a laser system in the femtosecond and picosecondranges. Void formation can be seen in the bulk of stent strut (below thesurface). FIG. 4B shows voids are dispersed through the surface regionof the cut edges.

It is believed that the high laser energy at the cut edge converts someof the inner solid mass into a volume of gas which results in theformation of voids or bubbles. The voids are dispersed within a regionextending from the cut edge to a given depth and beyond this depth thevoids dissipate. Such voids can act as stress concentrators that willfacilitates fracture formation when stent strut is under stress, andthus, reduce fracture resistance and reduce elongation at break. Inaddition to voids, cracking in the same region is also commonly observedin the samples.

FIG. 4A depicts a portion of a strut or structural element 500 formed bylaser machining a substrate. Strut 500 has an abluminal or outer surface(prior to machining the outer surface, e.g., of tube) and a lasermachined edge surface 508. The depth into the substrate normal to edgesurface 508 is shown by an arrow 510. FIG. 4B depicts a cross-section512 of a portion of strut 500 normal to machined edge surface 500 alongA-A. As shown by the cross-section, the substrate has voids or bubbles516 with a diameter Dv that extend to a depth Dp.

In addition, the modulus on the surface material of the stent (as afunction of distance from the machined edge surface for the stents) wasmeasured. Variation of modulus as a function of distance from themachined edge was observed. The variation becomes smaller as distanceincreases into the surface and tends to converge to the modulus of thevirgin polymer not affected by the laser energy as shown in the FIG. 9.Additional testing on the mechanical properties of the affected region,including ultimate strength, elongation at break, elastic modulus, andmaximum load were performed, and these mechanical outputs were adverselyeffected by the void formation.

The size of the voids and the depth to which they are present depend onlaser parameters such as wavelength and pulse width. The size of thevoids can be less than 1 micron, 1-2 microns, 2-5 microns, or greaterthan 5 microns. The cracks or voids may be present at no greater than 2microns, 5, 10, 15, 20, or 30 microns. In general, polymer stent'smechanical properties are affected by those voids and cracks formation,but it can be alleviated by proper choice of laser parameters such aswavelength and pulse width.

Additionally, thermal and nonthermal damage to a substrate materialdepends on the laser energy for a given pulse width for at least tworeasons. First, the optical penetration of the laser energy varies withwavelength. Second, the absorption coefficient, or more generally, thedegree of absorption of laser energy by a lased polymer varies withwavelength. In general, the lower the absorption coefficient, at a givenrange of wavelength, the greater is the thermal effect on the substrate.For example, the absorption coefficient of PLLA increases from anegligible value at 800 nm to reach a maximum at about 300-320 nm. Thus,photochemical removal increases and thermal removal decreases as thewavelength decreases from about 800 nm to about 300-320 nm. Thus, therelative amount of nonthermal or thermal ablation also depends on thewavelength of the laser because the polymer absorption coefficient isdependent on the wavelength of the laser.

Therefore, the relative amount of nonthermal/thermal ablation depends onboth the pulse width and the wavelength. Additionally, the degree ofphotochemical and thermal damage both depend on the pulse width.

As indicated above, the present invention includes adjusting parameters,such as pulse width and wavelength, to preserve or maintain mechanicalproperties of a stent and reduce or minimize damage to the uncut portionof the substrate. The adjusting can include selecting one or morewavelengths between ultraviolet (10-400 nm) and infrared (>700 nm) andone or more pulse widths. In some embodiments, a wavelength is selectedsuch that the polymer has an absorption coefficient at the wavelengththat is less than the maximum absorbance of the polymer, for example,the absorbance is 5-10%, 10-20%, 20-40%, or 40-60% of the maximumabsorbance.

In some embodiments, the pulse width is adjusted for the one or morewavelengths to avoid excessive melting, even with an appropriate levelof cooling during machining Excessive melting may correspond to greaterthan 0.25, 0.5, or greater than 1 micron thickness of melting material.

For a given pulse width and wavelength, the average laser power or power(energy per pulse×repetition rate) and repetition rate selected in orderto provide a fluence (energy per pulse/spot size of beam) that is highenough so that the beam cuts the substrate all the way through, forexample, the wall of polymer tubing. The beam spot size is generally10-20 microns, but can be less than 10 or greater than 20 micronsdepending on the application. A pulse energy and fluence (based on a 10micron spot size) for laser cutting polymers can be 4-200 μJ and 0.5-200J/cm², respectively. The average power per pulse of a beam can be 0.5-4W or more than 4 W. More narrowly, the power can be 0.5-1 W, 1-1.5 W,1.5-1.8 W, 1.8-2 W, 2-2.2 W, 2.2-2.5 W, 2.5-2.8 W, 2.8-3 W, 3-3.2 W,3.2-3.5 W, 3.5-3.8 W, 3.8-4 W. For a 10 ps pulse width laser, therepetition rate can be 25-100 kHz, 25-50 kHz, 50-60 kHz, 60-80 kHz, or80-100 kHz.

Additionally, the repetition rate and cooling gas flow rate (e.g., inSCFH He) are adjusted or selected in combination to reduce or minimizethermal effects (e.g., melting at surface of cut) and to maximizecutting speed. In general, the repetition rate and cutting speed aredirectly proportional, i.e., the faster the repetition rate, the fasterthe cutting speed, resulting in a lower process time per stent. However,as the repetition rate is increased, the thermal effects tend toincrease. An increase in the cooling gas flow rate can mitigate thethermal effects from the increased repetition rate, allowing a higherrepetition rate, and thus cutting speed. Thus, the repetition rate andcooling gas flow are chosen to obtain the fastest process time withacceptable thermal effects.

Although laser machining in a picosecond range, e.g., 1-12 ps or more,results in thermal effects, it may be advantageous to machine in thisrange since the thermal effects are more effectively controlled thanphotochemical effects. As indicated above, the thermal effects can bemitigated by cooling substrate as it is machined with a cooling gas.

Additionally or alternatively, minimizing the damage may correspond tominimizing the thickness of the affected region next to the lasermachined edge that includes voids. For example, the thickness of theregion can be between less than 2 microns, 5 microns, less than 20microns, or less than 30 microns. The void region can be 1-2 microns,2-5 microns, 2-10 microns, 2-20 microns, or 5-10 microns. Additionallyor alternatively, minimizing damage can help minimize the modulusvariation of the cut stent with distance in the damaged region. Theparameters may be adjusted to obtain the fasting convergence of themodulus toward a modulus of an undamaged polymer. The modulus mayconverge at less than 4 microns, less than 8 microns, or less than 20microns of the machined edge surface. The modulus may converge atbetween 1-4 microns, 4-8 microns, or 8-20 microns of the machined edgesurface. The modulus may converge at 4 microns or less, 8 microns orless, 15 microns or less, or 20 microns or less from the machined edgesurface.

Additionally or alternatively, damage may be minimized to the mostdesirable mechanical properties of the damaged region by adjusting thewavelength and pulse width, for example, the ultimate strength,elongation at break, elastic modulus, or maximum load. The mostdesirable is to have the cut polymer substrate maintaining the sameproperties prior to its laser machining.

The embodiments further include implementation a laser system andparameters in laser machining a PLLA substrate to manufacture a stent.In such embodiments, the laser wavelength may be in the visible lightspectrum from 390 to 800 nm. In some embodiments, the laser wavelengthis in the green spectrum or from about 496 to 570 nm, or more narrowly532 nm or 515 nm. In some embodiments, the pulse width can be 0.8 ps orless, 0.8-1 ps, 1-5 ps, 5-10 ps, 10-12 ps, 12-15 ps, or greater than 15ps. As disclosed in the examples, the inventors have demonstrated thatablating with a 10 ps laser with 532 nm wavelength at 80 kHz repetitionrate provides minimal damage to the PLLA stent as compared to othercombinations of wavelengths with lower pulse widths. In the greenwavelength range or at 532 nm specifically, the repetition rate of thelaser beam may be 25-100 kHz, or more narrowly, 25-40 kHz, 40-80 kHz, or80-100 kHz. Table 1 provides laser parameter ranges used in the study.

TABLE 1 Laser parameters Description Study range Wavelength, nm 532 nmPolarization circular Average laser power, W 1.5-3.0 Pulse energy  21 μJPulse width  12 ps Repetition rate  80 kHz for most stents 100 kHz forsome tests Gas (Helium) Flowrate, SCFH 6 to 10 Nozzle to substratedistance  0.9 mm Beam size  8 mm cutting speed, inch/min 4 to 16

The following definitions apply herein:

All ranges include the endpoints and any value within the endpoints,unless otherwise specified.

“Radial strength” of a stent is defined as the pressure at which a stentexperiences irrecoverable deformation.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. True stress denotes the stress where force and area aremeasured at the same time. Conventional stress, as applied to tensionand compression tests, is force divided by the original gauge length.

The “maximum load” or ultimate load is the absolute maximum load (force)that a structure can bear without failing.

“Strength” refers to the maximum stress along an axis which a materialwill withstand prior to fracture. The ultimate strength is calculatedfrom the maximum load applied during the test divided by the originalcross-sectional area.

“Modulus” may be defined as the ratio of a component of stress or forceper unit area applied to a material divided by the strain along an axisof applied force that results from the applied force. The modulus is theinitial slope of a stress-strain curve, and therefore, determined by thelinear Hookean region of the curve. For example, a material has both atensile and a compressive modulus.

“Strain” refers to the amount of elongation or compression that occursin a material at a given stress or load.

“Elongation” may be defined as the increase in length in a materialwhich occurs when subjected to stress. It is typically expressed as apercentage of the original length.

“Elongation to Break” is the strain on a sample when it breaks. It isusually is expressed as a percent.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. Tg of a givenpolymer can be dependent on the heating rate and can be influenced bythe thermal history of the polymer. Furthermore, the chemical structureof the polymer heavily influences the glass transition by affectingmobility.

EXAMPLES

The examples and experimental data set forth below are for illustrativepurposes only and are in no way meant to limit the invention. Thefollowing examples are given to aid in understanding the invention, butit is to be understood that the invention is not limited to theparticular materials or procedures of examples.

The following set of examples describes results of laser machiningstents from PLLA tubing for seven different parameter combinations ofpulse width and wavelength. The PLLA tubing was formed from an extrusionprocess from 100% PLLA resin. The dimensions of the tubing, the extrudeddimensions, are: outside diameter (OD)=0.0066 inch and inside diameter(ID)=0.0025 inch. The extruded PLLA tubes were radially expandedaccording to a process described previously, for example, in U.S.application Ser. No. 12/554,589, which is incorporated by referenceherein. The target percent radial expansion (% RE) was 400%, where % REis deformed as 100%×(Inside Diameter of Expanded Tube/Original InsideDiameter of Tube−1).

The seven different parameter combinations of pulse width and wavelengthare listed in Table 2.

TABLE 2 Pulse width and wavelength combination used to machine differentsets of substrates Set No. Pulse width Wavelength (nm) Laser 1   10 ps355 Trumpf¹ 2   10 ps 532 Lumera² 3 100 fs 266 Libra³ 4 100 fs 400Libra³ 5 100 fs 800 Libra³ 6 800 fs 1064 IMRA⁴ 7 800 fs 1550 Raydiance⁵¹Trumpf Laser Technology of Fremont, CA ²Lumera Laser GmbH ofKaiserslautern, Germany ³Coherent, Inc., Santa Clara, CA ⁴IMRA ofFremont, CA ⁵Raydiance of Petaluma, CA

Additional laser parameters for the runs are given in Table 3. The speedis the translational rate of the beam across the surface of thesubstrate. The passes is the number of times the cut pattern isrepeated. Processing time is the time required to cut an entire stent.The Coherent Libra (set #5) corresponds to the 800 nm in Table 3. ThePico (532 nm) lasers correspond to set #2. The lasers in Table 3 arefixed wavelength lasers.

TABLE 3 Summary of laser parameters Femto Pico (800 (355 Pico Pico PicoLaser Parameters nm) nm) (532 nm) (532 nm) (532 nm) Focusing Lens (mm)100 100 100 100 100 Spot Size (micron) 20 15 20 20 20 Rep Rate (kHz) 580 80 80 80 Power (mW) 160 200 1750 2100 2270 Speed (in/min) 4 10 4 8 16Passes 1 2 1 1 1 Helium (SCFH) 7 6-8 7 10 9 Processing time per 6:345:20 6:34 3:17 1:39 stent (min:sec)

The properties of the surface regions at the machined edge were assessedusing several testing techniques. Techniques and the properties obtainedfrom the tests are summarized in Table 4. Testing was performed on thesamples using cryo ultra-microtome, nanoindentation, and tensiletesting.

TABLE 4 Summary of test methods and properties studied Testing MethodProperties Cryo-ultra microtome/SEM Voids, cracks NanoindentationModulus vs. distance from cut surface Tensile testing Ultimate strengthElongation at break Elastic modulus Maximum load

The cryo-ultra microtome refers to a technique of cutting ultrathinsections for microscopic examination. Thin sections of the machinedstents normal to the machined edge were cut and examined with a scanningelectron microscope (SEM). The depth and size of voids and cracks forthe different sets were compared. Nanoindentation was used to measurethe modulus of the substrate as a function of distance from the machinededge. Tensile testing was used to measure the mechanical properties ofthe surface region from dog-bone structures cut from along the machinededge.

Cryo-Ultra Microtome Results

A summary of the results of cryo-ultra microtome technique for fourcombinations of pulse width and wavelength are given in Table 5 andFIGS. 5-8 show the SEM images of the surface region adjacent to themachined surface for each combination in Table 5. For set #5, apronounced photochemical effect is shown by bubbles extending to a depthof 20 μm. In set #4, the lower wavelength compared to set #5 appears tohave mitigated the photochemical effect since the depth of the bubblesis significantly less than set #5, however, the size of the bubbles inset #4 are larger. Contrary to what may have been expected, there is alarger depth for set #4 since the penetration depth and the wavelengthare inversely proportional.

TABLE 5 Summary of cryo-ultra microtome/SEM results Set No. Pulsewidth/wavelength Observation FIG. 5 100 fs/800 nm bubbles about 2 μmdiameter 5 20 μm depth 4 100 fs/400 nm bubbles about 1 μm diameter 6 5μm depth 1   10 ps/355 nm cracking and bubbles about 5 μm 7 diameter 30μm depth 2   10 ps/532 nm bubbles less than 1 μm diameter 8 2 μm depth

For set #1, cracks are present to a depth of 30 μm. Set #2 has the leastdamage among the four setups; the depth and the size of the bubbleformation are the lowest from all the test conditions.

Nanoindentation

A summary of the nanoindentation results with five combinations of pulsewidth and wavelength are given in Table 6 and FIGS. 9-13. The figuresshow the modulus vs. displacement into the surface of the machined edgefor 16 samples for each combination. Of the three combinations with apulse width of 100 fs, the 800 nm samples (set no. 5) show the fastestconvergence. Of the two combinations at 10 ps, the 532 nm wavelength hassignificantly faster convergence of the modulus. A comparison of the 100fs/800 nm and the 10 ps/532 nm results shows that the fastestconvergence of the modulus overall is provided by the 10 ps/532 nmwavelength combination.

TABLE 6 Summary of nanoindentation results Set No. Pulsewidth/wavelength Observation FIG. 5 100 fs/800 nm converges at about 8μm depth 9 4 100 fs/400 nm — 10 3 100 fs/266 nm — 11 1   10 ps/355 nmconverges at about 20 μm depth 12 2   10 ps/532 nm converges at about 4μm depth 13Tensile Testing

Tensile testing of surface regions of the machined edge was performedfor samples machined with pulse width/wavelength combinations of 100ps/800 nm and 10 ps/532 nm. The testing was performed using conventionaltensile testing equipment on dog-bone shape samples prepared fromsurface regions of laser machined substrates. Test details for the 100fs/800 nm sample and the 10 ps/532 sample are given in Tables 7 and 8,respectively.

TABLE 7 Description of tensile test sample using 100 fs/800 nm laserSamples Longitudinal dog-bones Rate 0.1 in/min Number of samples 19Thickness 0.0060 ± 0.0002 in. Width 0.0064 ± 0.0002 in.

TABLE 8 Description of the tensile test sample using 10 ps/532 nm laserSamples Longitudinal dog-bones Rate 0.1 in/min Number of samples 16Thickness 0.0061 ± 0.0001 in. Width 0.0060 ± 0.0001 in.

Table 9 summarizes the tensile test results. The most remarkableproperty difference is the elongation at break. The 10 ps/532 nm sampleshad an elongation more than 2.5 times or about 60% larger than the 100fs/800 nm samples. Therefore, the former samples have a significantlyhigher fracture resistance.

TABLE 9 Summary of tensile test results Percent Property 100 fs/800 nm10 ps/532 nm difference* Ultimate Strength (ksi) 12.5 ± 0.7 15.5 ± 0.919 Elongation (plastic) (%) 196 ± 66 496 ± 54 60 Elastic Modulus (ksi)58 ± 4 65 ± 3 11 Max. Load (lb_(f))  0.49 ± 0.03  0.57 ± 0.04 14 *100% ×(3rd col. property − 2nd col. property)/3rd col. propertySummary of Material Results

A summary of the material results for the sets of samples is provided inTable 10. The 10 ps/532 samples have the best properties.

TABLE 10 Summary of testing results. 100 ps/800 nm 10 ps/355 nm 10ps/532 nm Cryo-Ultra 20 um depth, 30 um depth, 2 um deep, <1 μmmicrotome 2 μm dia 5 μm dia dia Nano Converge @ Converge @ Converge @ 4μm Indentation 8 μm ~20 μm Tensile testing UTS 12.5 ksi, — UTS 15.5 ksi,496% 196% elongation elongationQuality of Machined Edge

The quality of the machined edge was investigated through observation ofSEM images of the sidewalls. One of the causes of roughness isredeposition of vaporized substrate material. FIGS. 14-17 depict SEMimages showing the sidewalls of stents machined with 100 fs/800 nm, 100fs/400 nm, 10 ps/355 nm, and 10 ps/532 nm, respectively. The sidewallmachined with the 10 ps/532 combination has the smoothest machined edge.

The above examples indicate that the 532 nm wavelength provides the bestbalance of cutting speeds with thermal effects and photochemicaleffects. This method can be applied to poly(L-lactide-co-glycolide) withany molar ratios of L-lactide to glycolide between 0 and 1. The lasersystem and its settings can also be applied to cut metallic substratessuch as Ta—Nb—W, and Co—Cr alloys.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A method of laser machining a substrate to form astent, comprising: providing a polymer substrate comprising a wall; andlaser machining the polymer substrate with a laser beam with a pulsewidth and wavelength that cuts through the wall to form structuralelements having a machined edge surface, wherein the laser beam modifiesthe substrate in a surface region adjacent to the machined edge surface,wherein the modifications include voids, cracks, variation in modulus ofthe polymer with distance from the edge surface, or a combinationthereof, wherein the wavelength of the laser beam is within the greenrange and the pulse width is 1-15 ps, and selecting the pulse width andwavelength so that the voids or cracks are present at no greater than adepth of 2 microns or the modulus converges at no greater than 4microns.
 2. The method of claim 1, wherein the polymer is PLLA or PLGA.3. The method of claim 1, wherein the polymer is PLLA and the wavelengthis 532 nm and the pulse width is less than or equal to 10 ps and greaterthan or equal to 1 ps.
 4. The method of claim 1, wherein the repetitionrate is 80-100 kHz.
 5. A method of laser machining a substrate to form astent, comprising: providing a polymer substrate comprising a wall; andlaser machining the polymer substrate with a laser beam to cut throughthe wall to form structural elements having a machined edge surface; andadjusting the pulse width and wavelength of the laser beam to minimizedamage to a surface region adjacent to the machined edge surface arisingfrom thermal and nonthermal ablation, wherein the wavelength of thelaser beam is within the green range and the pulse width is 1-15 ps,wherein the minimized damage comprises of cracks or voids no greaterthan a depth of 2 microns arising from nonthermal ablation and meltingno greater than 1 micron in thickness arising from thermal ablation. 6.The method of claim 5, wherein the polymer is PLLA and the pulse widthis adjusted to 1-10 ps.
 7. The method of claim 5, further directing acooling gas at a region of the substrate machined by the laser beam. 8.A method of laser machining a substrate to form a stent, comprising:providing a PLLA polymer substrate comprising a wall; and lasermachining the PLLA polymer substrate with a laser beam to cut throughthe wall to form structural elements having a machined edge surface,wherein the wavelength of the laser beam is within the green range andthe pulse width is 1-10 ps, wherein the wavelength in the green rangeand the pulse width are selected such that the laser beam causes cracksor voids to form in a surface region adjacent to the machined edgesurface, a depth of the voids or cracks being no greater than 5 microns.9. A method of laser machining a substrate to form a stent, comprising:providing a PLLA polymer substrate comprising a wall; and lasermachining the PLLA polymer substrate with a laser beam to cut throughthe wall to form structural elements having a machined edge surface,wherein the wavelength of the laser beam is within the green range andthe pulse width is 1-10 ps, wherein the laser beam causes variation inmodulus of the polymer with distance from the edge surface, wherein themodulus converges at a distance less than 10 microns.